Adaptive electromagnet for high performance magnetic resonance imaging

ABSTRACT

A method of configuring a conducting grid of elements interconnected at intersecting nodes by switches is described. The method includes: constructing a background grid by connection of centroids of the cell shape of the conducting grid; identifying a subset of elements in the background grid that intersect the smooth pattern of loops; identifying a subset of elements in the conducting grid that intersect the subset of elements in the background grid; the subset of elements in the conducting grid forming a discretized pattern of loops representing the smooth pattern of loops; for each of the discretized pattern of loops identifying current-in and current-out nodes; altering the on-off state of individual switches in accordance with the discretized pattern of loops; opening the switch between each respective pair of current-in and current out nodes; and applying power to the conducting grid via at least one pair of the input and output current nodes such that the current flow through the elements generates the magnetic field profile.

FIELD OF THE INVENTION

The present invention relates generally to magnetic resonance imaging.More specifically, the present invention relates to actively controllingthe spatial distribution of a magnetic field by use of an electromagnetwith an actively controllable current distribution.

BACKGROUND OF THE INVENTION

Magnetic resonance imaging (MRI) is a well known imaging technique thatcan be used to observe soft tissues such as the brain, muscles andkidneys. Specific properties of the various compounds found insidetissues, such as water and/or fat, are used to generate images. Whensubjected to a strong magnetic field, the vector sum of the nuclearmagnetic moments of a large number of atoms possessing a nuclear spinangular momentum, such as hydrogen, which is abundant in water and fat,will produce a net magnetic moment in alignment with the externallyapplied field. The resultant net magnetic moment can furthermore precesswith a well-defined frequency that is proportional to the appliedmagnetic field. After excitation by radio frequency (RF) pulses,relaxation mechanisms bring the net magnetization back to itsequilibrium position within a characteristic time T1 (also known as theT1 relaxation time), during which a signal can be detected. Theresulting MR image is a complex-valued map of the spatial distributionof the transverse magnetization M_(xy) in the sample at a specific timepoint after an excitation.

In MRI, the main magnetic field is produced by a large superconductingelectromagnet. Extreme care is taken to ensure that the magnetic fieldproduced by this magnet is uniform. Non-uniformities can result insignal loss, image distortion, image blurring, and poor fat suppression.In MR spectroscopy, field inhomogeneities cause broadening ofline-widths and frequency shifts. Due to these problems, great care istaken at the time of installation to ensure that the field produced bythe main magnet is extremely uniform; however, when a subject enters themagnetic environment, additional field inhomogeneities are produced dueto susceptibility differences between tissues. This problem is enhancedas the main magnetic field is increased. To achieve the stringent fielduniformity requirements necessary for MRI, both passive and activemagnetic shims are used to ‘fine-tune’ the main field in order to makeit as uniform as possible. See, for example, Romeo F., Hoult D. I.Magnet field profiling: analysis and correcting coil design, Magn ResonMed; 1: 44-65 (1984).

Typically, passive shims are utilized to remove inhomogeneities at thetime of installation and active room temperature electromagnets are usedto mitigate susceptibility induced field deviations. The active magneticshim coils traditionally consist of gradient coils (discussed in greaterdetail below) for first-order linear corrections, and an additional setof electromagnets that produce field patterns matching the second-orderspherical harmonics. Some high-field systems contain third oreven-fourth order shims. Each shim coil must be powered by its own powersupply, typically providing up to 10-20 A of current.

Spatial information in MRI, is encoded by linearly varying the mainmagnetic field using three room temperature electromagnets known asgradient coils. The gradient coils are typically located just inside the“bore” of the main magnet. The gradient coils produce magnetic fields onthe order of mT by passing hundreds of amperes of current through theirwindings. The power required to create these fields is provided byexpensive high-performance power amplifiers.

Due to heating and spatial constraints imposed by gradient coil designcriteria, the gradient coil fields can contain non-linearities as muchas 50% in extreme cases. The non-linearities result in image warping,which must be undone in post-processing of the image. The strength ofwhole-body gradient systems is in the range of 20-50 mT/m, withspecialized systems boasting strengths of 80-100 mT/m and dedicateddiffusion systems capable of 300 mT/m. Slew rates for the gradientsystems (i.e. how quickly they can be turned on) are around 200 T/m/s;however, due to the onset of peripheral nerve stimulation (PNS) mostscanners are operated at slew rates significantly lower than this.

Harris C. T., et al., A New Approach to Shimming: The DynamicallyControlled Adaptive Current Network, Magnetic Resonance Medicine, 71 pp.859-869 (2014), sets forth a dynamically controlled, activeelectromagnet that is capable of adaptively changing its wire patternfor the purpose of localized magnetic field shimming. Multiple differentspatial profiles can be produced (i.e. both the linear gradients andshim field patterns) using only a single electromagnet powered by asingle amplifier, thereby drastically reducing the cost and weightassociated with prior art systems. Furthermore, since the adaptiveelectromagnet can be positioned very close to the patient, lower poweris needed for a given field strength, eddy currents induced by switchingthe magnetic field are reduced if the system is further from the mainmagnet bore, field inhomogeneities with high spatial frequency can beaccounted for, and faster switching without the onset of PNS can beachieved.

A key requirement of the dynamic, adaptive electromagnet set forth inHarris et al. is the ability to represent a continuous current densitydistribution over a discretized grid of conducting material. However,Harris et al. does not provide any description of how a continuouscurrent density distribution can be transformed into a “discretized”pattern for application to a conducting grid, or any practicalimplementation of the dynamically controlled adaptive electromagnet.

Additional prior art is relevant to this specification:

-   Turner R., A target field approach to optimal coil design, J Phys D    Appl Phys; 19: L147-L151 (1986).-   Yoda K., Analytical design method of self-shielded planar coils. J    Appl Phys; 67: 4349-4353 (1990).-   Crozier S., Doddrell D. M. Gradient-Coil Design by Simulated    Annealing, J Magn Reson Ser A; 103: 354-357(1993).-   Lemdiasov R. A., Ludwig R. A Stream Function Method for Gradient    Coil Design, Concept Magn Reson B; 26B: 67-80 (2005).-   Poole M., Bowtell R. Novel gradient coils designed using the    boundary element method, Concept Magn Reson B; 33B: 220-227 (2007).-   Poole M., et al. Minimax current density coil design, J Phys D Appl    Phys; 43: 095001 (2010).-   Juchem C., et al. (2011). Multi-Coil Shimming of the Mouse Brain.    Magn Reson Med; 66: 893-900.-   Juchem C., et al. (2011). Dynamic Multi-Coil Shimming of the Human    Brain at 7 Tesla, J Magn Reson; 212: 280-288.-   Harris C. T., et al. Electromagnet design allowing explicit and    simultaneous control of minimum wire spacing and field uniformity,    Concept Magn Reson B; 41B(4): 120-129 (2012).

SUMMARY OF THE INVENTION

It is an object of an aspect of the invention to provide a novel systemand method for actively controlling the spatial distribution of amagnetic field in an MRI scanning system by use of an adaptiveelectromagnet, which obviates and mitigates at least one of theabove-identified disadvantages of the prior art.

According to one aspect, a method is set forth for transforming a smoothwire pattern to a discretized pattern that can be applied to aconducting grid. According to another aspect, a system and method areset forth for producing a desired current distribution in the conductinggrid according to the discretized pattern.

These, together with other aspects and advantages which will besubsequently apparent, reside in the details of construction andoperation as more fully hereinafter described and claimed, referencebeing had to the accompanying drawings forming a part hereof, whereinlike numerals refer to like parts throughout.

BRIEF DESCRIPTION OF THE DRAWINGS

FIG. 1 shows a block diagram of functional subsystems of a magneticresonance imaging (MRI) system in accordance with an implementation.

FIG. 2 shows an imaging volume and corresponding slice to be scanned bythe MRI system of FIG. 1 in accordance with an implementation.

FIG. 3 shows a simplified pulse sequence that may be used by the MRIsystem of FIG. 1 in accordance with an implementation.

FIG. 4 shows a schematic representation of a k-space data set containingone line received using the MRI system of FIG. 1 in accordance with animplementation.

FIG. 5 shows an exemplary rectangular conducting grid and a backgroundgrid offset by one half-grid spacing relative to the exemplaryrectangular conducting grid, in accordance with an embodiment;

FIG. 6 shows a triangular pattern of conducting grid and hexagonalpattern of background grid, in accordance with an embodiment;

FIG. 7 shows a hexagonal pattern of conducting grid and triangularpattern of background grid, in accordance with an embodiment;

FIG. 8 shows a mixed pattern of conducting and background grids, inaccordance with an embodiment;

FIG. 9 shows a portion of the background grid in FIG. 5 for generating adiscretized electromagnet wire pattern approximating a smooth wirepattern, according to an aspect of the invention;

FIG. 10 shows an exemplary rectangular background grid onto which thesmooth wire pattern has been superimposed;

FIG. 11 is a flowchart showing steps in an exemplary method fortransforming the smooth wire pattern to a discretized wire pattern onthe rectangular conducting grid, in accordance with an embodiment;

FIG. 12 shows elements in the background grid that intersect with thesmooth wire pattern, in accordance with an embodiment;

FIG. 13 shows elements of the rectangular conducting grid that intersectwith the background grid elements of FIG. 12, in accordance with anembodiment;

FIG. 14 shows elements of the rectangular conducting grid formingcurrent loops for generating a discretized wire pattern approximatingthe smooth wire pattern, according to an aspect of the invention;

FIG. 15 shows both the smooth and discretized patterns transformed froma two-dimensional representation to a cylinder;

FIG. 16 shows a detail of the lower-left portion of the rectangularconducting grid with conducting switches at each element;

FIGS. 17 and 18 comprise a flowchart showing steps for identifyingcurrent-in and current-out nodes for the current loops shown in FIGS. 12and 13, according to an aspect of the invention;

FIG. 19 shows the current-in and current-out nodes identified accordingto the method of FIGS. 17 and 18, according to an aspect of theinvention;

FIG. 20 shows a system for supplying power to a second conducting gridvia an intermediate third layer; and

FIGS. 21 and 22 show circuits for applying power to the rectangularconducting grid, according to an aspect of the invention.

FIG. 23 shows a system for supplying power to a second conducting gridvia an intermediate third layer where all switches are on the samesurface.

DETAILED DESCRIPTION

Referring to FIG. 1, a block diagram of a magnetic resonance imaging(MRI) system, in accordance with an example implementation, is shown at100. The example implementation of MRI system indicated at 100 is forillustrative purposes only, and variations including additional, fewerand/or varied components are possible.

As shown in FIG. 1, the illustrative MRI system 100 comprises a dataprocessing system 105. The data processing system 105 can generallyinclude one or more output devices such as a display, one or more inputdevices such as a keyboard and a mouse as well as one or more processorsconnected to a memory having volatile and persistent components. Thedata processing system 105 can further comprise one or more interfacesadapted for communication and data exchange with the hardware componentsof MRI system 100 used for performing a scan.

Continuing with FIG. 1, example MRI system 100 also includes a mainfield magnet 110. The main field magnet 110 can be implemented as apermanent, superconducting or a resistive magnet, for example. Othermagnet types, including hybrid magnets suitable for use in MRI system100 will now occur to a person of skill and are contemplated. Main fieldmagnet 110 is operable to produce a substantially uniform main magneticfield having a strength B0 and a direction along an axis. The mainmagnetic field is used to create an imaging volume within which desiredatomic nuclei, such as the protons in hydrogen within water and fat, ofan object are magnetically aligned in preparation for a scan. In someimplementations, as in this example implementation, a main field controlunit 115 in communication with data processing system 105 can be usedfor controlling the operation of main field magnet 110.

MRI system 100 further includes gradient coils 120 used for encodingspatial information in the main magnetic field along, for example, threeperpendicular gradient axes. The size and configuration of the gradientcoils 120 can be such that they produce a controlled and uniform lineargradient. For example, three paired orthogonal current-carrying primarycoils located within the main field magnet 110 can be designed toproduce desired linear-gradient magnetic fields.

In some implementations, gradient coils 120 may be shielded and includean outer layer of shield coils that can produce a magnetic field tocounter the gradient magnetic field produced by the primary gradientcoils forming a primary-shield coil pair. In such a coil pair the“primary” coils can be responsible for creating the gradient field andthe “shield” coils can be responsible for reducing the stray field ofthe primary coil outside a certain volume such as an imaging volume. Theprimary and shield coils of the gradient coils 120 may be connected inseries. It is also possible to have more than two layers of coils forany given gradient axis that together form a shielded gradient coil.Shielded gradient coils 120 may reduce eddy currents and otherinterference that can cause artifacts in the scanned images. Since eddycurrents mainly flow in conducting components of the MRI system 100 thatare caused by magnetic fields outside of the imaging volume (fringefields), reducing the fringe fields produced by gradient coils 120 mayreduce interference. Accordingly, the shapes and sizes, conductor wirepatterns and sizes, and current amplitudes and patterns of theprimary-shield coil pair can be selected so that the net magnetic fieldoutside the gradient coils 120 is as close to zero as possible. Forcylindrical magnets, for example, the two coils can be arranged in theform of concentric cylinders whereas for vertical field magnets, the twocoils may be arranged in coaxial disks.

The conductive components of the gradient coils 120, whether shielded orunshielded and including the primary and shield coils, may consist of anelectrical conductor (for example copper, aluminum, etc.). The internalelectrical connections can be such that when a voltage difference isapplied to the terminals of the gradient coils 120, electric current canflow in the desired path. The conductive components for the threegradient axes for both the primary gradient coils and the gradientshield coils can be insulated by physical separation and/or anon-conductive barrier.

The magnetic fields produced by the gradient coils 120, in combinationand/or sequentially, can be superimposed on the main magnetic field suchthat selective spatial excitation of objects occurs within the imagingvolume. In addition to allowing spatial excitation, the gradient coils120 may attach spatially specific frequency and phase information to theatomic nuclei placed within the imaging volume, allowing the resultantMR signal to be reconstructed into a useful image. A gradient coilcontrol unit 125 in communication with data processing system 105 isused to control the operation of gradient coils 120.

As discussed above, magnetic field “shims” may be used to improve theuniformity of the main magnetic field. To perform active shimming(correcting the field distortions that are introduced when differentobjects are placed within or around the system), correctiveelectromagnets, such as shim coils 139, carry a current that is used toprovide magnetic fields that act to make the main magnetic field moreuniform. For example, the fields produced by these coils may aid in thecorrection of inhomogeneities in the main magnetic field due toimperfections in the main magnet 110, or to the presence of externalferromagnetic objects, or due to susceptibility differences of materialswithin the imaging region, or any other static or time-varyingphenomena. A shim coil control unit 137 in communication with dataprocessing system 105 is used to control the operation of shim coils139.

Conventionally, magnetic shims fall into two categories: (1) passiveshims, composed of strategically placed ferromagnetic material withinthe magnet bore and/or superconducting electrical circuits within themagnet cryostat, and (2) active shims, composed of additionalroom-temperature electromagnets. Passive shims are typically used toadjust the main field at the time of initial installation, whereasactive shims are used to compensate for the field distortions that areintroduced when different objects are placed within the bore of themagnet.

Active shim coils are typically composed of sets of coaxial cylindricallayers, with each layer being a separate current path producing amagnetic field approximating a particular spherical harmonic. By drivingdifferent current amplitudes through each shim layer, the resultantadditive magnetic field profile can form complicated patterns. Thisapproach to active shimming can require significant amounts of radialspace, as each new spherical harmonic produced requires a newcylindrical coil. It also requires multiple power amplifiers, as eachcylindrical layer is driven separately. For higher performance, isdesirable to use a larger number of spherical harmonics, furtherincreasing radial space, power consumption, and number of amplifiersneeded.

As discussed in greater detail below with reference to FIGS. 5-23, amagnetic field shim is set forth in the form of an adaptableelectromagnet for actively controlling a magnetic field profile bydynamically adapting the pattern of current flow over a conducting grid.

Returning to FIG. 1, MRI system 100 further comprises radio frequency(RF) coils 130. The RF coils 130 are used to establish an RF magneticfield with strength B1 to excite the atomic nuclei or “spins”. The RFcoils 130 can also detect signals emitted from the “relaxing” spinswithin the object being imaged. Accordingly, the RF coils 130 can be inthe form of separate transmit and receive coils or a combined transmitand receive coil with a switching mechanism for switching betweentransmit and receive modes.

The RF coils 130 can be implemented as surface coils, which aretypically receive only coils and/or volume coils which can be receiveand transmit coils. RF coils 130 can be integrated in the main fieldmagnet 110 bore. Alternatively, RF coils 130 can be implemented incloser proximity to the object to be scanned, such as a head, and cantake a shape that approximates the shape of the object, such as aclose-fitting helmet. An RF coil control unit 135 in communication withdata processing system 100 can be used to control the operation of theRF coils 130.

There are many techniques for obtaining images using a MRI system 100,including T1 and T2 weighted images. To provide a simplifiedillustration of MRI system 100's functionality, simplified operationsfor obtaining proton density-weighted images are described as anon-limiting example. To create an image in accordance with the exampleillustration, MRI system 100 detects the presence of atomic nucleicontaining spin angular momentum in an object, such as those of hydrogenprotons in water or fat found in tissues, by subjecting the object to arelatively large magnetic field. In this example implementation, themain magnetic field has a strength B0 and the atomic nuclei containingspin angular momentum may be hydrogen protons. The main magnetic fieldpartially polarizes the hydrogen protons in an object placed in theimaging volume of the main magnet 110. The protons are then excited withappropriately tuned RF radiation, forming an RF magnetic field with astrength of B1, for example. Finally, weak RF radiation signal from theexcited protons is detected as an MR signal, as the protons “relax” fromthe magnetic interaction. The frequency of the detected MR signal isproportional to the magnetic field to which they are subjected.

Cross-sections of the object from which to obtain signals may beselected by producing a magnetic field gradient across the object sothat magnetic field values of the main magnetic field can be variedalong various locations in the object. Given that the signal frequencyis proportional to the varied magnetic field created, the variationsallow assigning a particular signal frequency and phase to a location inthe object. Accordingly, sufficient information can be found in theobtained MR signals to construct a map of the object in terms of protonpresence, which is the basis of a traditional MRI image. For example,since proton density varies with the type of tissue, tissue variationsmay be mapped as image contrast variations after the obtained signalsare processed.

Referring now to FIG. 2, to further illustrate the example signalacquisition process by the MRI system 100, it is presumed that an objectis placed within an imaging volume 250 of the main magnet 110 having amain magnetic field 210 with a strength B0, pointing along the Z-axisindicated at 240. The object subsequently has a net magnetizationvector. In this illustrative example, a slice in a plane along the X andY axes, as indicated at 205, is being imaged. It should be noted that inthis example, the slice has a finite thickness along the Z-axis,creating a volumetric slice 205.

To obtain images from the MRI system 100, one or more sets of RF pulsesand gradient waveforms (collectively called “pulse sequences”) areselected at the data processing system 105. The data processing system105 passes the selected pulse sequence information to the RF controlunit 135 and the gradient control unit 125, which collectively generatethe associated waveforms and timings for providing a sequence of pulsesto perform a scan.

Referring now to FIG. 3, an illustrative pulse sequence 300 is shownthat can be used to acquire images using the MRI system 100.Specifically, a timing diagram for the example pulse sequence isindicated at 300. The timing diagram shows pulse or signal magnitudes,as a function of time, for transmitted (RFt) signal, magnetic fieldgradients G_(x), G_(y), and G_(z) and a received RFx signal. The examplepulse sequence, simplified for illustrative purposes, contains a sliceselection radio frequency pulse 310 at RFt, a slice selection gradientpulse 320 at Gz, a phase encoding gradient pulse 330 at Gy, a frequencyencoding gradient pulse 340 at Gx, as well as a detected MR signal 350at RFx. The pulses for the three gradients Gx, Gy, and Gz represent themagnitude and duration of the magnetic field gradients that aregenerated by the gradient coils 120. The slice selection pulse 310 isgenerated by the transmit aspect of RF coils 130. Detected MR signal 350is detected by the receive aspect of the RF coils 130. In thisillustrative example it is presumed that transmit aspect and receiveaspect of RF coils 130 are formed by distinct coils.

The first event to occur in pulse sequence 300 is to turn on the sliceselection gradient pulse 320. The slice selection RF pulse 310 isapplied at the same time. In this illustrative example, the sliceselection RF pulse 310 can be a sinc function shaped burst of RF energy.In other implementations, other RF pulse shapes and durations can beused. Once the slice selection RF pulse 310 is turned off, the sliceselection gradient pulse 320 is also turned off and a phase encodinggradient pulse 330 is turned on. Once the phase encoding gradient pulse330 is turned off, a frequency encoding gradient pulse 340 is turned onand a detected MR signal 350 is recorded. It should be noted that theshapes, magnitudes and durations of the pulses and signals shown in FIG.3 are chosen for illustrative purposes, and that in implementations, oneor more of these factors and other signal factors may be varied toachieve the desired scan results.

In variations, the pulse sequence 300 can be repeated a certain numberof times or iterations, for example 256 times, to collect all the dataneeded to produce one image. Each repetition typically involvesvariations in the pulse sequence to allow receiving signalscorresponding to different components of the image. The time betweeneach repetition of the pulse sequence 300 can be referred to as therepetition time (TR). Moreover, the duration between the center point ofthe slice selection pulse 310 and the peak of detected MR signal 350 canbe referred to as echo time (TE). Both TR and TE can be varied asappropriate for a desired scan.

To further illustrate the signal acquisition process of MRI system 100,FIG. 2 is referred to in conjunction with FIG. 3. To select a slice, theslice selection gradient pulse 320 is applied along the Z-axis,satisfying the resonance condition for the protons located in the slice205. Indeed, the location of the slice along the Z-axis is determinedbased, in part, on the slice selective gradient pulse 320. Accordingly,the slice selection pulse 310, generated at the same time as the sliceselection gradient pulse 320 can excite protons that are located withinthe slice 205 in this example. Protons located above and below the slice205 are typically not affected by the slice selection pulse 310.

Continuing with the illustrative example, in accordance with the pulsesequence 300, a phase encoding gradient pulse 330 is applied after theslice selection gradient pulse 320. Since the gradient pulse 330 createsa gradient in the magnetic field along the Y-axis, the spins atdifferent locations along the Y-axis can begin to precess at differentLarmor frequencies. When the phase encoding gradient pulse 320 is turnedoff, the net magnetization vectors at different locations can precess atthe same rate, but possess different phases. The phases can bedetermined by the duration and magnitude of the phase encoding gradientpulse 330.

Once the phase encoding gradient pulse 330 is turned off, a frequencyencoding gradient pulse 340 can be turned on. In this example, thefrequency encoding gradient is in the X direction. The frequencyencoding gradient can cause protons in the selected slice to precess atrates dependent on their X location. Accordingly, different spatiallocations within the slice are now characterized by unique phase anglesand precessional frequencies. RF receive coils 130 can be used toreceive the detected signal 350 generated by the protons contained inthe object being scanned while the frequency encoding gradient pulse 340is turned on.

As the pulse sequence 300 is performed by MRI system 100, the acquiredsignals can be stored in a temporary matrix referred to as k-space, asshown in FIG. 4 at 400. Typically, K-space is the collection of thedetected signals measured for a scan and is in the spatial frequencydomain. K-space can be covered by frequency encoding data along theX-axis 420 (Kx) and phase encoding data along the Y-axis 430 (Ky)corresponding to the X and Y axis shown in FIG. 2. When all the linesfor the k-space matrix for a slice are received (at the end of the scanof a single slice, for example) the data can be mathematicallyprocessed, for example through a two-dimensional Fourier-transform, toproduce a final image. Thus, k-space can hold raw data beforereconstruction of the image into the spatial domain. K-space is filledwith raw data during the scan, typically one line per pulse sequence300. For example, the first line of k-space 400, indicated at 410, isfilled after the completion of the first iteration of the pulse sequencegenerated for scanning a slice and contains the detected signal for thatpulse sequence iteration. After multiple iterations of the pulsesequence, the k-space can be filled. Each iteration of the pulsesequence may be varied slightly, so that signals for the appropriateportions of the k-space are acquired. It should be noted that based ondifferent pulse sequences, other methods of filling the k-space arepossible, such as in a spiral manner, and are contemplated.

As discussed above, it is object of an aspect of the invention toprovide a novel system and method for actively controlling the spatialdistribution of a magnetic field in an MRI scanning system by use of anadaptive electromagnet. According to an embodiment, by modifying currentpathways as a function of time, multiple different magnetic fieldprofiles are created with one or more surfaces driven by one or morepower supplies. Furthermore, by modifying the current pathways during apulse sequence, the field profile may be dynamically altered over alocalized region of interest (ROI), on a patient-specific basis, in realtime.

As discussed in greater detail below, according to an exemplaryembodiment conducting switches are placed at the nodes of a conductinggrid to adaptively control the current flowing through the grid. Byaltering the on-off state of the switches, the current flow alongdifferent pathways may be controlled, giving rise to multiple differentmagnetic field profiles.

In one embodiment, a set of pre-defined current paths can be stored inthe shim coil control unit 137, which can then switch between thedifferent pre-defined current paths in a time-varying manner. In thisembodiment multiple sets of either those switches that are to beactivated, or switches that are to be deactivated, or both, can bestored in shim coil control unit 137.

In another embodiment, a current path (or wire pattern) is created insuch a way so as to produce a desired magnetic field profile. This wirepattern can be created by a variety of methods, as set forth for examplein Turner R.: A target field approach to optimal coil design, J Phys DAppl Phys; 19: L147-L151, Crozier S., Doddrell D. M.: Gradient-CoilDesign by Simulated Annealing, J Magn Reson Ser A; 103: 354-357,Lemdiasov R. A., Ludwig R.: A Stream Function Method for Gradient CoilDesign, Concept Magn Reson B; 26B: 67-80, or Poole M., Bowtell R.: Novelgradient coils designed using the boundary element method, Concept MagnReson B; 33B: 220-227. All of these approaches result in a smooth wirepattern, that is, the wires can flow anywhere over the surface asopposed to a pattern that has been restricted over a pre-determinedconducting grid. The next step in this embodiment is to transform thesmooth wire pattern to one that can be represented on a discretizedconducting grid 500. FIGS. 5-19, and the accompanying description below,set forth a computationally efficient, robust, method for discretizing asmooth wire pattern for application to the conducting grid.

The exemplary embodiment discussed herein utilizes a conducting gridwith a second “background grid”. The two grids can be different shapes,as set forth below. In the following description, the term “cell shape”describes the shape of the grid (e.g. square, rectangle, hexagon,triangle, etc.); “elements” are the sides of the cell shape; and “nodes”are the vertices of the cell shape. With this notation, the backgroundgrid can be described as the set of lines that join together thecentroids of each cell shape on the conducting grid. Examplerepresentations of conducting and background grids are shown in FIG. 5for a rectangular conducting grid 500 and rectangular background grid900; FIG. 6 for a triangular conducting grid 500 and hexagonalbackground grid 900; FIG. 7 for a triangular background grid 900 andhexagonal conducting grid 500, and lastly, FIG. 8 for mixed shapeconducting and background grids 500 and 900, respectively.

The elements of the conducting grid can be active (conducting) orinactive (non-conducting). The method set forth below is used toidentify which elements should be active that will best represent thesmooth wire pattern. With reference to FIG. 9, two current pathways areshown 510 representing portions of a smooth wire pattern. This patternis superimposed on top of a rectangular background grid at 900. Thesmooth wire pattern has a minimum separation distance between twocurrent pathways denoted 6. In order for the smooth pattern to berepresented by a conducting grid 500, discussed further below, theminimum separation distance must be larger than the maximum distancebetween the nodes of the background grid 900. In the illustrated examplethe maximum distance between the nodes of the rectangular backgroundgrid is the diagonal, denoted D. For non-rectangular grids, such asdepicted in FIGS. 6-8, this distance is shown with a double-sided arrow,also denoted D. It is important to note that for mixed cell shape grids(FIG. 8), the maximum distance between the nodes of the background grid900 (and hence the limit for the minimum separation between two currentpathways on the smooth wire pattern) will vary depending on the densityand shape of the grid. Therefore, the grid can be discretized morefinely in certain areas that require a smoother representation of theoriginal wire pattern or where the minimum separation between twocurrent pathways must be small.

To achieve the required minimum separation between conductors imposed bythe grid discretization, the number of loops can be reduced in the wirepattern representation of the current density, or a design method can beused that distinctly incorporates a minimum wire separation distanceinto its optimization (see Poole M., et al.: Minimax current densitycoil design, J Phys D Appl Phys; 43: 095001 or Harris C. T., et al.:Electromagnet design allowing explicit and simultaneous control ofminimum wire spacing and field uniformity, Concept Magn Reson B; 41B(4):120-129).

FIG. 10 displays a smooth wire pattern 510 that is to be represented bya rectangular conducting grid. The smooth wire pattern is shownsuper-imposed on top of the background grid 900 described previously.FIG. 11 shows steps for representing the smooth wire pattern by theconducting grid. At step 1100, the smooth wire pattern is super-imposedonto the coordinate system of the conducting grid. In the illustratedexample this step entails transforming from the 3D Cartesian coordinatesof the smooth wire pattern to Polar coordinates, where the azimuthalangle θ is the local x surface-coordinate and z is the local ysurface-coordinate. According to this exemplary embodiment, theconducting grid is a 2 cm×2 cm square grid of conducting pathways in thexy-plane (in surface coordinates, or θz-plane in polar coordinates)spanning a total area of 94 cm×60 cm in the x- and y-directions,respectively. This corresponds to a cylindrical surface with radius ofapproximately 14.95 cm and total length of 60 cm that has been“unwrapped” so that the planar x-direction corresponds to the azimuthaldirection in polar coordinates and the planar y-direction corresponds tothe z-direction in polar coordinates.

At step 1110, the background grid 900 is constructed. For this exemplarycase the background grid has the same discretization size as theconducting grid 500 (i.e. 2 cm×2 cm grid), and offset by exactly onehalf-grid spacing (1 cm) in both the x- and y-directions, as shown inFIG. 5. Next, at step 1120, wire elements 1000 in the background grid900 that intersect with the smooth pattern 510 are identified, as shownin FIG. 12. Then, at 1130, once the highlighted elements 1000 (that is,elements in the background grid that intersect the smooth wire pattern)have been identified, the elements 1100 of the conducting grid 500 thatintersect with the highlighted elements 1000 of the background grid areidentified, as shown in FIG. 13. The elements 1100 form a set of gridcurrent loops 1200-1255, etc., that best represent the smooth wirepattern, as shown in FIG. 14, wherein the grid current loops aresuper-imposed on the smooth wire pattern representation 510. FIG. 15shows both the smooth and discretized patterns transformed from theunwrapped plane in polar coordinates to a cylinder in Cartesiancoordinates. FIG. 16 shows a detail of the lower-left portion of grid500 with switches at each element.

Care must be taken to ensure that the specified current direction forthe discretized grid pattern 1100 is the same as the smooth wire pattern510 by, for example, performing a dot product between the intersectingelement from the smooth wire pattern and the representative element inthe grid.

The method of FIG. 11 maintains separation between loops (i.e. there isno connection from loop to loop). Depending on how the electromagnet isdriven, connections between loops may or may not be needed. In somecircumstances it may be desirable to have connections between loops,which can be accomplished many ways. In one example implementation thiscan be done by identifying “current-in” and “current-out” nodes and thenconnecting the current-in and current-out nodes, as set forth in FIGS.17 and 18.

At step 1700, an initial starting node 1510 for the “current-in” (i.e.the current from the amplifier) is specified. As shown in FIG. 19, thestarting current input node 1510 is chosen to be the bottom left-mostnode of the grid 500. At 1710, the loop that is nearest to the initialcurrent-in node 1510 in the x-direction is identified. In general therecan be a series of loops that have the same distance to the startingnode in the x-direction (e.g. loops 1200-1225), when this occurs onemust select the starting loop based on secondary criteria (e.g. closestin the y-direction). Therefore, in this example, the initial loop willbe 1200. Next, at step 1720, a determination is made as to whether anyof the remaining loops fall within the initial loop, and, if so,proceeds to step 1730 ordering them according to their distance from theinitial current-in node by the previously described criteria (e.g. theorder would be 1205, 1210). Then, if no further current loops remainwithin the initial loop 1200 (i.e. a NO branch at step 1720), the loop1215 that is next nearest to the initial current-in node 1510 isidentified (step 1735). Successive loops 1220-1225 are likewiseidentified (steps 1720-1740) until no further loops remain that areequidistant from the starting node 1510 in the x-direction. Then (step1760), the process repeats at 1710, identifying the next set of loopsthat are nearest to the starting node, excluding the loops that havealready been ordered until there are no further loops. In this examplethe next set of loops will be ordered as 1230, 1235, 1240, 1245, 1250,1255, etc.

Steps for identifying current input and current output nodes for eachloop are set forth in FIG. 18. At 1800, the element of the first loop ofthe previously ordered set that is nearest to the starting node 1510 isidentified (e.g. in the illustrated embodiment this is the element thatjoins nodes 1510 and 1500), and selecting one of the nodes at the end ofthis element as the current in and, at 1810, one as the current outdepending on the direction of current flow through the loop. In thisexample, the current in node is 1510 and the current out node is 1500.This process is repeated for the successive closest loops until allloops have current-in and current-out nodes (1520-1550, etc.) associatedwith them (steps 1820-1840). The final current-out node is the node thatis connected to the power amplifier (e.g. in the far right corner of thegrid 500). FIG. 19 displays the current-in and current-out nodes for thefirst twelve loops of the exemplary grid shown in FIGS. 14 and 15. Itshould be noted that after the current-in and current-out nodes for aparticular loop have been identified, the element in the loop that joinsthe nodes together is removed (i.e. the switch on the element isopened), as shown in FIG. 16.

Having described a method for creating a discretized wire pattern aswell as an algorithm to select the current in and current out nodes foreach loop, a method and circuit for supplying power to the conductinggrid is now set forth for producing the desired current distribution.

In one hardware implementation, a single power-in node and a singlepower-out node are provided for connecting the power supply to theconducting grid 500. In one embodiment, depicted in FIG. 20, theconnections between loops are made via a second grid layer 2000, whileconnection between the first and second grid layers is made via a thirdlayer 2010 placed between the first and second layers for separating allnodes of the first and second layers other than nodes that correspond toa “current-in” or “current-out” node as described above. This third,separation layer 2010, can have a switch located between the nodes ofthe two conducting layers, wherein the switch is closed if the nodecorresponds to a current in or current out node and is otherwise open,as shown in FIG. 20. It should be noted that placement of the singlepower-in and power-out nodes can be arbitrarily set, and their positionscan be varied depending on the desired current pattern.

An alternate embodiment to FIG. 20 is illustrated in FIG. 23 which showsa system for supplying power to a second conducting grid via anintermediate third layer where all switches are on the same surface. Onebenefit of this implementation is that it saves on space. For example,if all switches are on the top surface, it would allow for the backside(bottom surface) of a multilayer conductor to be used to route thecontrol signals for the switches.

Although the foregoing implementation requires only a single amplifier,the grid network may be split into multiple areas each with a singlepower-in and power-out node with a corresponding power supply therebyallowing for additional flexibility in producing field profiles.

According to another embodiment, each loop may be assigned a current-inand current-out node independently of the other loops in the desiredpattern. In this implementation connections between loops are not neededthereby reducing the number of required grid surfaces. According to thisembodiment, each element of the conducting grid 500 may operate as acombined current-in and current-out line for a given loop. Thus, for agiven loop 1200, the element joining nodes 1500 and 1510 serves as thecurrent-in line to the loop, the switches on the conducting grid 500cause current to flow around the loop back to this element and then flowback down a separate pathway out of the grid to the return terminal ofthe power supply. The switch placed on this element (i.e. the switchseparating nodes 1500 and 1510 is open, thereby restricting current toflow directly from the current-in line to the current-out line withoutfirst flowing around the loop. Each loop 1200-1255, etc. requires onesuch combined current-in/current-out node and a separate power supply(or a circuit to equally divide the current source from a single powersupply to each loop), but connections between loops are not required. Ina current dividing implementation it should be noted that the amount ofcurrent delivered to each loop need not be the same (i.e. one loop canreceive 1 A and the next loop can receive 2 A), permitting comparableperformance to a “multi-coil” approach with greater flexibility over thechoice of basis current loop shape.

One example of an implementation of the foregoing embodiment is shown inFIGS. 21 and 22. A square conducting grid 500 is provided on one surfaceof a cylinder, as set forth above (see FIG. 15). The grid containsswitches on the elements between its nodes, which control the currentpathway over that grid. The current pathway is identified by the methodsdescribed previously. Two strips of conducting material are distributedin the azimuthal direction and oriented in the z-direction on a secondsurface 1720 at the same spacing as the conducting grid on the firstsurface but offset half the grid spacing so as to be centered on theswitches connecting the nodes of the grid. One end of the conductingstrips is connected to a power supply (i.e. one strip connected to thepositive terminal and the other connected to the negative terminal) andthe other end is unconnected to create an open circuit. At eachz-position where an element occurs (i.e. where a switch is placed on theconducting grid 500) there are two vertical vias on either side of theswitch (FIG. 22) that allow current to flow from the two conductingstrips on layer 1720 to the conducting grid 500 on the underlying layer.Within these vias (or on an intermediary layer between the two layers)there is a switch on each line that can either be open or closed. Beforethe current pathway is determined, all switches on the intermediarylayer are open. Once a current loop has been determined, a currentin/out element is selected on the loop and the switches for that elementare closed thereby allowing current to flow from the power supply to oneside of that element, through the loop to the other side of the elementand back to the power supply. For simplicity, one can restrict theopening and closing of the switches so that only one element is open foreach longitudinal pair of conducting strips. The above-describedembodiments are intended to be examples and alterations andmodifications may be effected thereto, by those of skill in the art,without departing from the scope which is defined solely by the claimsappended hereto. For example, methods, systems and embodiments discussedcan be varied and combined, in full or in part.

We claim:
 1. A method of configuring a conducting grid of elementsinterconnected at intersecting nodes by switches, for generating adynamically changeable current distribution represented in twodimensions on said conducting grid by a smooth pattern of loops, tocreate a time varying magnetic field profile within a region of interestin an MRI system, comprising: constructing a background grid to saidconducting grid by connection of centroids of the cell shape of theconducting grid; identifying a subset of elements in said backgroundgrid that intersect said smooth pattern of loops; identifying a subsetof elements in said conducting grid that intersect said subset ofelements in the background grid, wherein said subset of elements in saidconducting grid forms a discretized pattern of loops representing saidsmooth pattern of loops; for each of said discretized pattern of loopsidentifying input and output current nodes; altering the on-off state ofindividual switches in accordance with said discretized pattern ofloops; opening the switch between each respective pair of input andoutput current nodes; and applying power to said conducting grid via atleast one pair of said input and output current nodes such that thecurrent flow through said elements generates said magnetic fieldprofile.
 2. The method of claim 1, further comprising interconnectingrespective ones of said discretized pattern of loops via said input andoutput current nodes.
 3. The method of claim 1, wherein identifyinginput and output current nodes further comprises: identifying an initialcurrent-in node for receiving said power; ordering said discretizedpattern of loops based on distance from said initial current-in node;and identifying a current-in and current-out node for each successiveloop in said ordering.
 4. An adaptive electromagnet, comprising: anarray of conducting grid elements for generating a time varying magneticfield profile within a region of interest in an MRI system; a pluralityof switches interconnecting said conducting grid of elements atintersecting nodes for generating a dynamically changeable currentdistribution represented in two dimensions on said array of conductinggrid elements by a smooth pattern of loops; a background gridconstructed to said array of conducting grid elements by connection ofcentroids of the cell shape of the array of conducting grid elements,wherein a subset of elements in said background grid intersect saidsmooth pattern of loops and a subset of elements in said array ofconducting grid elements intersect said subset of elements in thebackground grid such that said subset of elements in said conductinggrid forms a discretized pattern of loops representing said smoothpattern of loops; and input and output current nodes for each of saiddiscretized pattern of loops; wherein upon (i) altering the on-off stateof individual switches in accordance with said discretized pattern ofloops, (ii) opening the switch between each respective pair of input andoutput current nodes, and (iii) applying power to said conducting gridvia at least one pair of said input and output current nodes, currentflows through said elements to generate said magnetic field profile, andwherein said switches are transistor switches.
 5. The adaptiveelectromagnet of claim 4, wherein said elements are arranged in arectangular grid comprising a plurality of horizontal and vertical wireconductors that intersect at said plurality of nodes.
 6. The adaptiveelectromagnet of claim 4, wherein the input input current node of atleast one loop is connected to a corresponding output current node of anext loop via a second conductive grid.
 7. The method of claim 1,wherein applying power to said array further comprises: assigning eachloop a combined current-in and current-out element; and connecting aseparate source of power to each respective combined current-in andcurrent-out element.
 8. An adaptive electromagnet, comprising: an arrayof conducting grid elements for generating a time varying magnetic fieldprofile within a region of interest in an MRI system; a plurality ofswitches interconnecting said conducting grid of elements atintersecting nodes for generating a dynamically changeable currentdistribution represented in two dimensions on said array of conductinggrid elements by a smooth pattern of loops; a background gridconstructed to said array of conducting grid elements by connection ofcentroids of the cell shape of the array of conducting grid elements,wherein a subset of elements in said background grid intersect saidsmooth pattern of loops and a subset of elements in said array ofconducting grid elements intersect said subset of elements in thebackground grid such that said subset of elements in said conductinggrid forms a discretized pattern of loops representing said smoothpattern of loops; a combined current-in and current-out element for eachof said discretized pattern of loops; a separate source of power to eachrespective combined current-in and current-out element; wherein upon (i)altering the on-off state of individual switches in accordance with saiddiscretized pattern of loops, (ii) opening the switch between eachrespective combined current-in and current-out element, and (iii)applying power to said conducting grid via one said respective combinedcurrent-in and current-out element, current flows through said gridelements to generate said magnetic field profile, and, wherein thesource of power comprises a power supply.
 9. An adaptive electromagnet,comprising: an array of conducting grid elements for generating a timevarying magnetic field profile within a region of interest in an MRIsystem; a plurality of switches interconnecting said conducting grid ofelements at intersecting nodes for generating a dynamically changeablecurrent distribution represented in two dimensions on said array ofconducting grid elements by a smooth pattern of loops; a background gridconstructed to said array of conducting grid elements by connection ofcentroids of the cell shape of the array of conducting grid elements,wherein a subset of elements in said background grid intersect saidsmooth pattern of loops and a subset of elements in said array ofconducting grid elements intersect said subset of elements in thebackground grid such that said subset of elements in said conductinggrid forms a discretized pattern of loops representing said smoothpattern of loops; a combined current-in and current-out element for eachof said discretized pattern of loops; a separate source of power to eachrespective combined current-in and current-out element; wherein upon (i)altering the on-off state of individual switches in accordance with saiddiscretized pattern of loops, (ii) opening the switch between eachrespective combined current-in and current-out element, and (iii)applying power to said conducting grid via one said respective combinedcurrent-in and current-out element, current flows through said gridelements to generate said magnetic field profile, and wherein the sourceof power comprises a single power supply and a circuit to divide thecurrent from the power supply to each loop via each combined current-inand current-out element.
 10. The adaptive electromagnet of claim 9,wherein the amount of current to each loop is equal.
 11. The adaptiveelectromagnet of claim 9, wherein the amount of current to each loop isunequal.
 12. The adaptive electromagnet of claim 8, further comprising:first and second strips of conducting material distributed in theazimuthal direction on a surface adjacent said conductive grid, saidstrips overlying switches of the conductive grid, one strip beingconnected at one end to a positive terminal of said power supply and theother strip being connected at said one end to a negative terminal ofsaid power supply and an opposite end of each strip left unconnected tocreate an open circuit; a pair of vias at each azimuthal positionoverlying opposite sides of a switch of said conductive grid forallowing current to flow from the first and second strips of conductingmaterial to said grid array; a further switch associated with each viaadapted to be closed in response to assigning each loop acurrent-in/current-out element and altering the on-off state ofindividual switches on said grid array in accordance with saiddiscretized pattern of loops, thereby allowing current to flow from thepower supply to one side of the current-in/current-out element, througheach said loop and to the other side of the current-in/current-outelement then back to the power supply.
 13. The adaptive electromagnet ofclaim 12, wherein said further switch is disposed within said via. 14.The adaptive electromagnet of claim 12, wherein said further switch isdisposed in an intermediary layer between said surface and saidconductive grid.
 15. The adaptive electromagnet of claim 9, furthercomprising: first and second strips of conducting material distributedin the azimuthal direction on a surface adjacent said conductive grid,said strips overlying switches of the conductive grid, one strip beingconnected at one end to a positive terminal of said power supply and theother strip being connected at said one end to a negative terminal ofsaid power supply and an opposite end of each strip left unconnected tocreate an open circuit; a pair of vias at each azimuthal positionoverlying opposite sides of a switch of said conductive grid forallowing current to flow from the first and second strips of conductingmaterial to said grid array; a further switch associated with each viaadapted to be closed in response to assigning each loop acurrent-in/current-out element and altering the on-off state ofindividual transistor switches on said grid array in accordance withsaid discretized pattern of loops, thereby allowing current to flow fromthe power supply to one side of the current-in node/current out element,through each said loop and to the other side of thecurrent-in/current-out element then back to the power supply.
 16. Theadaptive electromagnet of claim 15, wherein said further switch isdisposed within said via.
 17. The adaptive electromagnet of claim 15,wherein said further switch is disposed in an intermediary layer betweensaid surface and said conductive grid.
 18. The adaptive electromagnet ofclaim 16, wherein all switches are located on the same surface.